Non-Uniformly Stiff Polymeric Scaffolds and Methods for Producing Thereof

ABSTRACT

The invention relates to methods for producing a polymeric scaffold for use in tissue engineering applications or soft tissue surgery, as well as to the produced scaffolds and an associated kit. The method features a first fast drying step of applying a mechanical compression on a polymeric gel layer and a second slow drying step of the gel up to reach a polymer mass fraction of at least 60% w/w in the final scaffold. The method allows the production of scaffolds with high regeneration and healing properties of a grafted tissue via host cell invasion and colonization, and a good suturability. These goals are achieved through the formation within the scaffold of a non-uniform architecture creating softer and stiffer areas, which is maintained even upon re-swelling of the scaffold upon hydration of the final dried product.

TECHNICAL FIELD

The invention relates to tissue engineering, diagnostics and surgery,particularly to non-uniformly stiff polymeric scaffolds and methods forproducing such objects.

BACKGROUND ART

Collagen is one of the first natural polymers used as a biomaterial.Being the most abundant structural extracellular matrix protein in thehuman body, it is therefore widely used as a scaffold component fortissue engineering applications, and considered among the first choicesfor natural scaffold materials.

Collagen gels are formed by a process called fibrillogenesis.Fibrillogenesis starts when a diluted acidic collagen solution isadjusted in vitro to physiological pH and temperature (Gross and Kirk,J. Bio. Chem 233 (1958): 355-360). Collagen gels are interesting asscaffolds due to their good biocompatibility, their low immunogenicityand their easy of remodeling by cells. However, the large excess offluid entrapped between the collagen fibers render them inherently weakand thus difficult to manipulate. Several strategies were proposed toobtain collagen scaffolds having improved mechanical properties. Thesimple culturing of cells within collagen gels results in the expulsionof some of the interstitial fluid by gel contraction, but cellcompaction is a not an easily controllable process and the increase incollagen density is linked to a reduction of the scaffold surface (i.e.scaffold shrinkage). Cross-linking of the collagen fibers enhances themechanical properties of collagen scaffolds without changing thescaffold surface. Physical, chemical and biological collagencross-linking methods are well reported in literature (Paul and Bailey,The Scientific World Journal 3 (2003): 138-155). One option is thelyophilization (freeze-drying) of an aqueous collagen solution followedby chemical cross-linking (EP1561480). Another possibility isfreeze-drying of a collagen gel followed by thermal dehydrationcross-linking (EP1098024).

The most common aldehyde cross-linking agent is glutaraldehyde. Howeverits use for medical applications is questioned due to its cytotoxicity.Several chemical cross-linking agents have been identified that achievecomparable cross-linking levels to that of glutaraldehyde whileimproving the cytocompatibility (i.e. diphenylphosphoryl azide,1-ethyl-3-(3-dimethylaminopropyl)carbodiimide, genipin) (Petite H. etal., Biomaterials 16 (1995): 1003-1008).

Another method to produce collagen gels with enhanced mechanicalproperties consists of applying plastic compression. During the processof plastic compression, the interstitial liquid is expelled from acollagen gel and does not return on removal of the load. This simple,fast, controllable, and cell-friendly process to engineer collagenmatrices with suitable mechanical properties for soft tissue engineeringwas described in detail by Brown and colleagues (WO 2006/003442; WO2012/004564; Abou Neel E. A. et al., Soft Matter 2 (2006): 986-992;Brown R. A. et al., Adv Funct Mater 15 (2005): 1762-1770). Furthermore,collagen hybrid structures have been reported where a synthetic polymercomponent provided the mechanical support (Kawazoe N. et al., BiotechnolProg 26 (2010): 819-826; Lu H. et al., Sci Technol Adv Mater 13 (2012)).However, this approach decreases the biocompatibility of the matrix.

Most scaffold design strategies aim to mimic the mechanical propertiesof the native tissue as much as possible while also providing a scaffoldthat is easy to handle and suturable. The following suturable collagenmatrices were reported in literature: collagen sponge orcollagen/glycosaminoglycan sponge crosslinked with collagen gel (U.S.Pat. No. 5,567,806), double layered collagen matrix resulting from acasting, freeze-drying, and lyophilisation procedure where the collagenlayers differ in density (WO 2007/082295), and a high densityfreeze-dried collagen matrix (WO 2000/029484).

Furthermore it is well known that the mechanical properties of ascaffold influences cell behaviour. Myogenesis is best achieved oncollagen scaffolds that are moderately stiff (Engler A. J. et al., JCell Biology 166 (2004): 877-887) while osteogenesis is achieved onstiff collagen scaffolds (Huebsch N. et al., Nat Matter 9 (2010):518-526). Researchers can change mechanical properties of collagenhydrogels by varying collagen concentration and using differentcross-linking methods. However, these scaffolds are uniform matrices anda soft collagen scaffold showing excellent cell behaviour in term ofproliferation and differentiation might not be suitable for a clinicalapplication (i.e. not suturable, difficult to handle).

SUMMARY OF INVENTION

In order to address and overcome the drawbacks associated with the priorart solutions concerning polymeric products for tissue engineering andsurgery, the inventors developed a brand new method for producingpolymeric scaffolds displaying several advantages when used in an invivo approach. Particularly, the scaffold of the present inventionfeatures superior properties in terms of resistance upon suture whengrafted in vivo in/on soft tissues, while exhibiting at the same timeexcellent mechanical and functional properties in terms of facilitationof host cell infiltration upon grafting, reduced inflammatoryreactions/rejection and improved tissue healing, with associated fasterrecovery of a subject undergone to a surgical operation. The scaffold ofthe invention can be stored off-the-shelf, thus reducing the costsassociated to the surgery, and providing successful grafting resultseven in a cell-free scenario. In a particular aspect, the presentinvention proposes an easy and controllable process to providenon-uniformly stiff polymeric scaffolds that consists of denser areasensuring suturability and easy handling for the surgeons, as well assofter zones to induce the appropriate cell behaviour and fate in termof cell proliferation and differentiation.

Accordingly, it is an object of the present invention to provide amethod for producing a non-uniformly stiff polymeric scaffold for use intissue engineering, diagnostics or surgical procedures, characterized inthat it comprises the following steps:

-   -   a) providing at least one layer of a polymeric material gel;    -   b) performing a first fast drying step by applying a mechanical        compression on said polymeric material gel layer; and    -   c) performing a second slow drying step of the gel up to reach a        polymeric material mass fraction of at least 60% w/w.

In one embodiment, the method is followed by a hydration step where thescaffold re-swells with more than 4% w/w.

In one embodiment, the method is characterized in that step c) isperformed by water filtration means, air-drying or any combinationthereof.

In a particular embodiment, the method is characterized in that waterfiltration is performed by dialysis.

In a particular embodiment, the method is characterized in thatair-drying is performed under gravity for at least 1 hour.

In one embodiment, the method further comprises the step ofcircumferentially wrapping the polymeric scaffold around a support inorder to have a tubular polymeric scaffold comprising a lumen.

In one embodiment, the method comprises step a′) instead of step a) ofpouring a polymeric gel precursor solution into a substantiallycylindrical mould comprising therein a coaxial elongated support forcreating a tubular single layer of polymeric material gel.

In one embodiment, the method is characterized in that the polymericscaffold is cell-free.

In one embodiment, the method is characterized in that the polymericmaterial is natural polymeric material or an extra-cellular matrix(ECM)-derived polymeric material.

In one embodiment, the method is characterized in that the naturalpolymeric material or ECM-derived polymeric material is gelatin,elastin, collagen, agar/agarose, chitosan, fibrin, proteoglycans, apolyamino-acid or its derivatives, preferably polylysin or gelatinmethyl cellulose, carbomethyl cellulose, polysaccharides and theirderivatives, preferably glycosaminoglycanes such as hyaluronic acid,chondroitinsulfate, dermatansulfate, heparansulfate, heparine,keratansulfate or alginate, or any combination of the foregoing.

In one embodiment, the method is characterized in that no crosslinkersare added to the polymeric material gel.

In a particular embodiment, the method is characterized in that thepolymeric gel of step a) is obtained from a collagen gel precursorsolution having a collagen concentration of at least 2.1 mg/mL,preferably between 2.5 and 40 mg/mL.

A further aspect of the invention relates to a method for producing anon-uniformly stiff polymeric scaffold for use in tissue engineering,diagnostics or surgical procedures, characterized in that it comprisesthe following steps:

-   -   a) providing a single layer of a mechanically compressed        polymeric material gel wrapped around a support;    -   b) dipping a portion of the support-wrapped polymeric gel layer        into an aqueous solution while leaving another portion        air-drying under gravity for a certain amount of time;    -   c) flipping the polymeric gel layer portion previously dipped        into the aqueous solution onto the air-dried portion; and    -   d) optionally air-drying under gravity both layers for a certain        amount of time        wherein at least one layer of the so obtained polymeric scaffold        has a polymer mass fraction of at least 60% w/w.

In one embodiment, said method comprises step a′) instead of step a) ofpouring a polymeric gel precursor solution into a substantiallycylindrical mould comprising therein a coaxial elongated support forcreating a tubular single layer of polymeric gel, gelling it andmechanically compress it.

In one embodiment, the method is characterized in that the polymericscaffold is cell-free.

In one embodiment, the method is characterized in that the polymericmaterial is natural polymeric material or an extra-cellular matrix(ECM)-derived polymeric material.

In one embodiment, the method is characterized in that the naturalpolymeric material or ECM-derived polymeric material is gelatin,elastin, collagen, agar/agarose, chitosan, fibrin, proteoglycans, apolyamino-acid or its derivatives, preferably polylysin or gelatinmethyl cellulose, carbomethyl cellulose, polysaccharides and theirderivatives, preferably glycosaminoglycanes such as hyaluronic acid,chondroitinsulfate, dermatansulfate, heparansulfate, heparine,keratansulfate or alginate, or any combination of the foregoing.

In one embodiment, the method is characterized in that no crosslinkersare added to the polymeric material gel.

In a particular embodiment, the method is characterized in that thepolymeric gel of step a) is obtained from a collagen gel precursorsolution having a collagen concentration of at least 2.1 mg/mL,preferably between 2.5 and 40 mg/mL.

A further aspect of the present invention relates to a non-uniformlystiff polymeric scaffold for use in tissue engineering, diagnostics orsurgical procedures produced according to the previously-describedmethods, said scaffold being characterized in that it has at least apolymeric material layer having a polymer mass fraction of at least 52%w/w at steady state upon re-swelling, with a denser polymeric structureat the surface and less dense in the core of the scaffold.

In one embodiment, the non-uniformly stiff polymeric scaffold ischaracterized in that it is suturable.

In one embodiment, the non-uniformly stiff polymeric scaffold ischaracterized in that it is at least 1 cm long.

In one embodiment, the non-uniformly stiff polymeric scaffold ischaracterized in that it comprises a bioactive agent.

Another object of the present invention relates to a kit comprising:

-   -   a) a container;    -   b) an aqueous solution; and    -   c) the non-uniformly stiff polymeric scaffold as previously        described.

In one embodiment, the kit is characterized in that is sterilizedthrough electrode beam, gamma-rays radiation or X-rays radiation.

BRIEF DESCRIPTION OF DRAWINGS

FIG. 1 depicts a sketch of the method of the invention for theproduction of sheets (1A) or tubes (1B) of a non-uniformly stiffpolymeric scaffold;

FIG. 2 shows one embodiment of a mold used to produce a polymerictubular scaffold with an inner lumen;

FIG. 3 shows a schematic representation of the “sock technique”according to the invention for producing non-uniformly stiff polymericscaffolds;

FIG. 4 shows images of the “sock technique” and the final structure of adouble-layered and a single-layered collagen tube, as well as a collagensheet obtained by opening of a double-layered tube. (A) The upper halfof a collagen tube is hydrated. (B-E) The hydrated collagen half ispulled over the dry half. (F) A single layered suture compatiblecollagen tube on a forceps. (G) A double-layered suture compatible tubeon a forceps. (H) The opened up double-layered tube forming a doublelayered sheet;

FIG. 5 shows the influence of the drying method and the final dryingstate on the re-swelling behaviour of collagen scaffolds. (A) Mechanicalcompressed followed by air-drying is compared to only air-drying ofcollagen, with monitoring of the water content and time of air-drying.There is a clear trend in the time of air-drying and water content (B)No drying of produced collagen scaffolds results in a 100% water contentin the collagen scaffold (white bars after drying), and the watercontent does not change upon incubation in a hydrating environment(black bars after re-swelling). Utilizing the method of freeze-drying toreduce water content in the collagen scaffold results in a driedscaffold with an average of 1.3% water content, and reaches an averageof around 20% water content after re-swelling. Utilizing only air-dryingallows obtaining an average dry level of 1.5% water content thatincreases to 9.5% after re-swelling. Utilizing mechanical compressionwith air-drying allows obtaining an average of 0.8% water content thatincreases to 5.2% after re-swelling. Utilizing only mechanicalcompression allows obtaining an average dry level of 47.6% water contentthat increases to 48.3% after re-swelling. (C) The air-drying method canbe used to manipulate the re-swelling behaviour of collagen scaffolds bychanging the final dry-state (% water); A linear trend between the finaldry-state of the collagen scaffold and the degree of re-swelling isobserved; At least 1 h of air drying after mechanical compression yieldsa different material with different water content compared to only amechanical compressed scaffold;

FIG. 6 shows the re-swelling steady state of collagen scaffolds. Freezedried, air-dried, and mechanically compressed with air-dried collagenscaffolds rapidly re-swell (within 30 min) in a hydrating environment. Asteady-state re-swelling is occurring within 72 h in hydratingconditions;

FIG. 7 shows the relation between the mechanical properties andpercentage of water present within surgery-compatible collagenscaffolds. (A and B) Air-dried collagen scaffolds show that waterpercentage is a crucial parameter for the mechanical properties (Youngmodulus and Ultimate tensile strength) of the scaffold. (C and D) Anair-dried (1.5% water content) and re-swelled collagen scaffold (9.5%water content) is anastomosed to a goat urethra using Vicryl 6.0sutures. (E) Even a mechanical compressed plus air-dried for 1 h (38%water content) and re-swelled collagen scaffold (45% water content) isanastomosed to a dog urethra using Vicryl 6.0 suture. (F and G)Hematoxylin & eosin staining of tissue samples harvested 1 month and 3months after surgery in rabbits. White arrows show the sutures and blackarrows point to the remaining collagen of the implanted scaffold that isslowly being replaced by host tissue and finally completely disappeared3 months post-surgery;

FIG. 8 shows the microscopic collagen structure at the micro-nano scaleafter different drying methods. A and D show a collagen tube produced bymechanical compression (plastic compression) followed by air drying(semi-elastic compression) yielding a heterogeneous or “non-uniform”collagen structure. The non-uniform structure is the result of theapplication of two different drying methods during the fabricationprocedure. B and E shows a collagen tube that has only been air-driedyielding a homogenous or “uniform” collagen structure. C and F show acollagen tube that has been freeze-dried yielding a random, non-uniformcollagen structure, with pronounced holes distributed randomly in thecollagen network;

FIG. 9 visualizes the different collagen areas present within a tubularscaffold that was produced by applying both compressions, the plastic(mechanical load) and the semi-elastic (air-drying) one. (A) Schematicdescription of the manipulations to obtain a non-uniformly stiff tubularcollagen scaffold. (B) Schematic representation of the collagenstructure of the final tubular scaffold. (C and D) Immunohistochemistryusing anti-collagen type 1 antibodies on samples of a tubular bovinecollagen scaffold to visualize the collagen type 1 fiber density. Imagestaken under a LSM700 confocal microscope (Zeiss) visualize the surfaceof the inner lumen (D) and the outer side (C) of a cross-section of acollagen tube;

FIG. 10 shows the morphology of urothelial cells cultured on collagenscaffolds with different densities. GFP-lentivirial transducedurothelial cells were used to easily visualize them under a fluorescencestereomicroscope (Leica). 24 h after cell seeding, urothelial cellsgrowing on (A) less dense collagen scaffolds form more single units andless connected colonies as compared to (B) urothelial cells growing onhigher density collagen gels.

DESCRIPTION OF EMBODIMENTS

The present disclosure may be more readily understood by reference tothe following detailed description presented in connection with theaccompanying drawing figures, which form a part of this disclosure. Itis to be understood that this disclosure is not limited to the specificconditions or parameters described and/or shown herein, and that theterminology used herein is for the purpose of describing particularembodiments by way of example only and is not intended to be limiting ofthe claimed disclosure.

As used herein and in the appended claims, the singular forms “a”, “an”and “the” include plural referents unless the context clearly dictatesotherwise. Thus, for example, reference to “a bioactive agent” includesa plurality of such agents and reference to “a layer” includes referenceto one or more layers, and so forth.

Also, the use of “or” means “and/or” unless otherwise stated. Similarly,“comprise”, “comprises”, “comprising”, “include”, “includes” and“including” are interchangeable and not intended to be limiting. It isto be further understood that where descriptions of various embodimentsuse the term “comprising”, those skilled in the art would understandthat in some specific instances, an embodiment can be alternativelydescribed using language “consisting essentially of” or “consisting of”.

The main object of the present invention relies at least in part on anovel method for producing a polymeric material scaffold for use intissue engineering or surgical procedures, as well as for diagnosticsand basic research, having very peculiar mechanical and functionalcharacteristics. In fact, once produced with the methods according tothe invention, the scaffolds of the invention surprisingly show anon-uniform architecture, wherein some zones within the polymericscaffold display differences in polymer density and therefore also instiffness. Since there is a close relation between the polymerconcentration and the mechanical properties of the scaffold such asstiffness and elasticity, this feature advantageously guide cellmigration, proliferation and differentiation in vivo thanks to thepresence of one or more less dense polymeric area that can be easilyremodeled by invading host cells, thus enhancing the tissue regenerationcapacity of the scaffold, while keeping at the same time a portion ofthe scaffold with a sufficient mechanical strength for being suturedduring surgery.

The method of the invention covers two main approaches for themanufacturing of the non-uniformly stiff polymeric scaffolds of theinvention. Contrary to some prior art approaches involving a process ofplastic compression for creating uniform polymer (e.g. collagen)matrices, the present method exploits a new method renamed by theinventors “semi-elastic compression” of polymeric gels. In contrast toan elastic compression, where the polymeric gel swells back to theirinitial state after removal of the compressive force applied therein,the gels processed according to the present method swell back only to anintermediate state, which is dependent on the initial polymerconcentration and the final dry state. Surprisingly, these semi-elasticcompressed and rehydrated polymeric structures are stable enough to beeasily handled and to be sutured in an in vivo surgical approach. Thisis realized by applying a first fast mechanical compression step usingmechanical load followed by a slow drying phase, preferably followed bya final rehydration.

The invention will be better understood with the help of the followingdefinitions.

As used herein, a “polymeric material” is any material comprisingpolymers, large molecules (also known as macromolecules) composed ofmany repeated smaller units, or subunits, called monomers, tightlybonded together preferably by covalent bonds. Polymer architecture atthe molecular scale can be rather diverse. A linear polymer consists ofa long linear chain of monomers. A branched polymer comprises a longbackbone chain with several short side-chain branches covalentlyattached. Cross-linked polymers have monomers of one long or short chaincovalently bonded with monomers of another short or long chain.Cross-linking results in a three-dimensional molecular network; thewhole polymer is a giant macromolecule. Another useful classification ofpolymers is based on the chemical type of the monomers: homopolymersconsist of monomers of the same type, copolymers have differentrepeating units. Furthermore, depending on the arrangement of the typesof monomers in the polymer chain, there are the followingclassification: the different repeating units are distributed randomly(random copolymer) or there are alternating sequences of the differentmonomers (alternating copolymers) in block copolymers long sequences ofone monomer type are followed by long sequences of another type; andgraft copolymers consist of a chain made from one type of monomer withbranches of another type. A sufficiently dense polymer solution can becrosslinked to form a polymer gel, including a hydrogel or a cryogel,which is a soft solid.

Polymer materials may also be formed by blending two or more polymersinto physical mixtures. For example, the poor impact strength of apolymer can be greatly improved by incorporating small particles of anelastomer. Many properties of polymeric materials depend on themicroscopic arrangement of their molecules. Polymers can have anamorphous (disordered) or semicrystalline (partially crystalline,partially ordered) structure. Polymers can be mixed with inorganicparticles (usually in the form of continuous fibres, such as glass orparticulates such as mica, talc and clay) in order to modify and improve(mainly but not exclusively) their mechanical properties.

Suitable polymeric materials include, but are not limited to, syntheticpolymers such as polyurethanes, poly-olefins, polyethylene glycol (PEG),poly(glycolide) (PGA), poly(L-lactide) (PLA) orpoly(lactide-co-glycolide) (PLGA); natural polymeric materials (i.e.,non-synthetic polymers, polymers that can be found in nature) and/orpolymers derived from the extracellular matrix (ECM) as gelatin,laminin, elastin, collagen, agar/agarose, chitosan, fibrin,proteoglycans, a polyamino-acid or its derivatives, preferably polylysinor gelatin methyl cellulose, carboxymethylcellulose (CMC),polysaccharides and their derivatives, preferably glycosaminoglycanessuch as hyaluronic acid, chondroitinsulfate, dermatansulfate,heparansulfate, heparine, keratansulfate or alginate, nucleotides,polylipides, fatty acids, as well as any derivative thereof, fragmentthereof and any combination of the foregoing. Natural and ECM-derivedpolymers are a first choice biomaterial for tissue engineeringapplications envisaged by the present disclosure, due to theirbiological and chemical similarities to natural tissues and the presenceof biologically active sites in their structures. For example, collagenand hyaluronic acid can be considered as the most preferred embodimentsfor constituting the polymeric scaffolds of the invention.

As used herein, the term “gel” refers to a non-fluid colloidal networkor polymer network that is expanded throughout its whole volume by afluid. A gel is a solid three-dimensional network that spans the volumeof a liquid medium and ensnares it through surface tension effects. Theinternal network structure may result from physical bonds (physicalgels) or chemical bonds (chemical gels).

As used herein, the term “hydrogel” refers to a gel in which theswelling agent is water. A hydrogel is a macromolecular polymer gelconstructed of a network of crosslinked polymer chains. It issynthesized from hydrophilic monomers, sometimes found as a colloidalgel in which water is the dispersion medium. Hydrogels are highlyabsorbent (they can contain over 90% water) natural or syntheticpolymeric networks. As a result of their characteristics, hydrogelsdevelop typical firm yet elastic mechanical properties.

Several physical properties of the (hydro)gels are dependent uponconcentration. Increase in (hydro)gel concentration may change its poreradius, morphology, or its permeability to different molecular weightproteins. One skilled in the art will appreciate that the volume ordimensions (length, width, and thickness) of a (hydro)gel can beselected based on instant needs, such as for instance the region orenvironment into which a scaffold substantially made of an (hydro)gel isto be implanted.

For “collagen” is herein meant any kind of natural or recombinantcollagen of animal origin, preferably mammals, such as bovine, porcine,rat, human and the like. Accordingly, collagen type 1 to 28 can be usedin the frame of the present invention, as well as fragments thereof orany combinations of the foregoing. Collagen can also possibly derivefrom any body location of the said animals such as skin, intestine,cartilage and so forth, as long it fits with biocompatibilityrequirements or with translation into an in vivo approach. In apreferred embodiment, type 1 collagen of bovine origin is used.

The polymeric gel of the invention can be manufactured/processed throughany suitable manufacturing method known in the art allowing to create ahighly interconnected pore network in the material, such as e.g.(photo)lithography including two-photon lithography, casting andmolding, 3D printing, inkjet printing, porogen leaching, emulsionfreezing/freeze drying technique, inverse opal hydrogelation,cryogelation, electrospinning or fiber extrusion and bonding, gasfoaming, thermally induced phase separation and so forth.

In one embodiment, the polymeric gel can be fabricated by top-downprocesses such as 3D printing. This technique allows full control overthe distribution and size of pores and wall material in nearly arbitrarygeometries, and therefore holds great possibilities in terms oforgan-specific 3D design and fabrication.

In the frame of the present disclosure, a “scaffold” is any threedimensional material having a framework architecture, i.e. a supportstructure usually comprising hollow spaces within it. Generallyspeaking, a scaffold material is an artificial, usually temporarytemplate structure capable of supporting three-dimensional bodytissue/organ formation in vivo, ex vivo or in vitro. In this context, ascaffold material is also referred herewith as a “biomaterial” or“bioscaffold”. A bioscaffold, inter alia, allows cell attachment andmigration, delivers and retains cells and biochemical factors, enablesdiffusion of vital cell nutrients and expressed products, exerts certainmechanical and biological influences to modify the behaviour of thecells and so forth.

The scaffold material of the invention has been conceived andmanufactured in order to act as a biocompatible, non-migratory andnon-carcinogenic scaffolding agent. The general purpose was thedevelopment of a material having improved long-term efficacy, which canstimulate host cell infiltration and integrates with the surroundingtissue once implanted in a host, thus triggering neo-tissue formationvia the promotion of a bioactive environment within the application siteand giving rise to long-term functional repair. The scaffold materialherein disclosed addresses and solves, among others, this problem.

The scaffold material is moreover, at least on some portions thereof,highly porous. In a preferred embodiment of the invention, the pores areall interconnected in order to create a continuous net of material thatcan act as a plausible physical support for other elements such as cellsor bioactive agents, while providing at the same time additional keyfeatures to the scaffold such as its softness/stiffness, easiness ofcell/tissue invasion and so forth. Accordingly, where present, porediameters preferably comprised between about 0.2 and 200 μm in the lessdense areas, with more preferred pore diameters comprised between about0.5 and 10 μm. As will be evident to a person skilled in the art, theseless dense areas within the scaffold are present, due to the non-uniformstiffness (and therefore density) of the scaffold itself, in the softer(and therefore, less dense) portion, while the stiffer (and therefore,denser) portion thereof is characterized by a negligible porosity, withpores, where present, with an average diameter of less than 0.2 μm.Moreover, it is preferred that the stiffer portion of the polymericscaffold has a Young's modulus comprised between about 5 and 1500 kPa,whereas preferred Young's modulus values for the softer portion spanbetween about 0.1 and about 20 kPa. In any embodiment, however, at leasta 2.5 kPa difference in terms of Young's modulus between the softest andstiffest part of the scaffold is present.

In order to optimize the mechanical properties of the scaffold materialof the invention and, in some aspects, its resorption/biodegradationrate, in preferred embodiments it is contemplated that the scaffoldcomprises a final polymer weight in a dried form of at least the 60% ofthe total scaffold weight (i.e. 60% w/w, also called mass fraction,wherein the remaining 40% or less is substantially composed of water orof an aqueous solution). A suitable polymer mass fraction depends one.g. the molecular weight of the monomer, the nature of the monomer, thecrosslinking strategy and the ratio of polymers. However, this value istightly associated to the manufacturing steps of the method forproducing said polymeric scaffold; in fact, as it will be more deeplydetailed later on, the final polymer mass fraction in the scaffold caneven be of at least 52% w/w upon re-swelling (re-hydration) of thescaffold in an aqueous solution, at the end of the production procedure.

Concerning the possible degradation/resorption rate of the scaffold uponin vivo application/implantation in a subject, this is mainly dependenton physico-chemical properties of the polymeric material of which it iscomposed of, as well as further factors such as crosslinking of thepolymers, the polymer concentration, the site of implantation into ahost and the like. The degradation/resorption rate can be calibrated byadjusting said physico-chemical parameters, such as for instance bypolymer crosslinking, the use of inhibitor molecules, by changing thepolymer density, crystallinity and/or its molecular weight distribution,changing the materials' porosity and so forth. Generally speaking, thescaffold material may be, at least in part and at least in some portionthereof, intrinsically biodegradable in vivo.

The scaffolds and the gels of the invention are preferably cell-free,but they can be seeded in vitro or in vivo with cells before, duringand/or after their production. The scaffold, if to be used to transplantcells in a tissue engineering approach, comprises pores to permit thestructure to be seeded with cells and e.g. to allow the cells toproliferate. For example, scaffolds are seeded by incubating thestructure in a solution containing the cells for a suitable period oftime. Alternatively, a multi-component scaffold is built in stages witheach layer being seeded prior to laying down of another layer or beforeadherences of another pre-formed component. Alternatively, eachcompartment of the multi-compartment scaffold can be seeded separately.Different cell types, e.g., stem vs. differentiated, and/or with variousphenotypes in terms of differentiation, activation, metabolic orfunctional state, can be seeded in the scaffold material. The scaffoldsare suitable for use with any cell type, preferably forming softtissues, which one may want to transplant. Such cells include but arenot limited to, various stem cell populations (embryonic stem cellsdifferentiated into various cell types), bone marrow or adipose tissuederived adult stem cells, mesenchymal stem cells, cardiac stem cells,pancreatic stem cells, neuronal cells, glial cells, spermatozoids andoocytes, endothelial progenitor cells, outgrowth endothelial cells,dendritic cells, hematopoietic stem cells, neural stem cells, satellitecells, side population cells. Such cells may further include but are notlimited to, differentiated cell populations including osteoprogenitorsand osteoblasts, chondrocytes, keratinocytes for skin, intestinalepithelial cells, smooth muscle cells, cardiac muscle cells, epithelialcells, endothelial cells, urothelial cells, fibroblasts, myoblasts,chondroblasts, osteoclasts, hepatocytes, bile duct cells, pancreaticislet cells, thyroid, parathyroid, adrenal, hypothalamic, pituitary,ovarian, testicular, salivary gland cells, adipocytes and combinationsthereof. For example, smooth muscle cells and endothelial cells may beemployed for muscular or tubular scaffolds, e.g., scaffolds intended asvascular, oesophageal, intestinal, rectal, or ureteral scaffolds;chondrocytes may be employed in cartilaginous scaffolds; cardiac musclecells may be employed in heart scaffolds; hepatocytes and bile ductcells may be employed in liver scaffolds; myoblasts may be used inmuscle regeneration; epithelial, endothelial, fibroblast, and nervecells may be employed in scaffolds intended to function as replacementsor enhancements for any of the wide variety of tissue types that containthese cells. In general, scaffolds of the invention may comprise anycell population competent to participate in regeneration, replacement orrepair of a target tissue or organ, particularly (semi)soft orbarely-reachable ones such as kidneys, brain, meninges, lungs, liver,stomach, sphincters, gut, bladder, ear tissues, cartilages, skin,pharynx, trachea, blood vessels, urethra, ureter, vagina, pelvic flooror pancreas.

Some mechanical/functional properties of the scaffold can be tailoredaccording to the needs by changing the physical or chemical propertiesthereof (such as e.g. the collagen molecular chain length). Concerningcollagen, in order to optimize these parameters and, in some aspects,the resorption/biodegradation rate, in preferred embodiments it iscontemplated an average molecular weight for the collagen moleculessubstantially composing the scaffold comprised between about 30 andabout 250 kDa. Concerning hyaluronic acid, an average molecular weightcomprised between about 50 kDa and about 2 MDa is particularlypreferred.

In most preferred embodiments, the final scaffold is not crosslinked orminimally crosslinked, in order to eliminating or reducing as much aspossible e.g. the potential in vivo toxicity associated with somecrosslinking agents such as glutaraldehyde. When a crosslinking processis used, crosslinking agents and their amount can be chosen at theoperator's discretion, and a skilled in the art would easily envisagesuch parameters based on common practice. By practicing the productionmethod of the invention that will be detailed later on, it is possibleto produce scaffolds with such mechanical and functional properties toallow at the same time successful grafts in a host with easesuturability, without the impellent need of using crosslinking agents.For being considered successful, a graft should show no complications atthe grafting site, such as no presence of opened structures as fistulae,strictures or stenoses, and/or the treated tissue/organ area should lookhistologically normal.

In some embodiments, the scaffold can be differentially permeable,allowing cell displacement only in certain physical areas thereof. Thepermeability of the scaffold composition is regulated, for example, byengineering the polymeric material for greater or smaller pore size,porosity, density, polymer cross-linking and/or viscoelasticity.

The scaffold material of the invention can be organized, during or afterthe production method, in a variety of geometric shapes such as beads,pellets, strip, block, toroid, capillary, patches, tubes, planar layers(e.g., thin sheets) and so forth, the final shape depending on theparticular context it is to be used, and can comprise in someembodiments at least two areas or compartments having differentstructural/functional properties. For example, multicomponent scaffoldsare constructed in e.g. concentric layers, each of which ischaracterized by different physico-chemical properties (% polymer, %crosslinking of polymer, chemical composition of scaffold, pore sizeand/or architecture, stiffness, porosity, presence or differentconcentration of bioactive agents and so on). Each niche can host one ormore cell populations and/or have a specific effect on them, e.g.,promoting or inhibiting a specific cellular function, proliferation,differentiation, migration and so forth. Such a configuration isparticularly useful in maintaining for long time periods the “sternness”of a population of cells, while simultaneously pushing daughter cells tomultiply rapidly and differentiate appropriately for participation intissue regeneration.

A compartmentalization of the device can be obtained by rational designand fabrication processes using different compositions or concentrationsof compositions for each compartment. For example, a stem cellpopulation can be encapsulated within hydrogels, using standardencapsulation techniques. Alternatively or additionally, two differentareas or compartments can be formed within the bioscaffold that containsdistinct factors (e.g., morphogens, growth factors, adhesion ligands),or the material in a distinct form (e.g., varying mechanical propertiesor porosity). In some embodiments, the compartments can be fabricatedindividually, and then operably coupled to each other with any suitablemethod known in the art such as use of an adhesive or exploitation ofthe intrinsic adhesiveness of each compartment, polymerization orcross-linking of one compartment to another and so forth. The scaffoldcan be designed to have a number of compartments or areas in which cellsenter in parallel and distribute according to their characteristics(e.g., size or mobility), serially pass through all or some of thecompartments, or a combination of both. The different compartments caneven be construed to induce distinct fates for the contained cellsduring the passage therein.

During or after the manufacturing process, the scaffold material can befunctionalized with additional elements such as for instance bioactivemolecules. Said elements can be coated on or embedded within thebioscaffold with any suitable means known in the art, either in ahomogeneous or non homogenous manner and can provide additionalfunctional properties to the material such as enhanced/reducedbiodegradation, physical stabilization, biological activity and thelike. As used herein, a “bioactive molecule”, as well as “(bio)activecompound” or “therapeutic agent”, is any active agent having an effectupon a living organism, tissue, or cell. The expression is used hereinto refer to any compound or entity that alters, inhibits, activates, orotherwise affects biological or chemical events.

One skilled in the art will appreciate that a variety of bioactivecompounds can be used depending upon the needs. Exemplary therapeuticagents include, but are not limited to, a small molecule, a growthfactor, a protein, a peptide, an enzyme, an antibody or any derivativethereof (such as e.g. multivalent antibodies, multispecific antibodies,scFvs, bivalent or trivalent scFvs, triabodies, minibodies, nanobodies,diabodies etc.), an antigen, a nucleic acid sequence (e.g., DNA or RNA),a hormone, an anti-inflammatory agent, an anti-viral agent, ananti-bacterial agent, a cytokine, an oncogene, a tumor suppressor, atransmembrane receptor, a protein receptor, a serum protein, an adhesionmolecule, a lypidic molecule, a neurotransmitter, a morphogeneticprotein, a differentiation factor, an analgesic, organic molecules,metal particles, radioactive agents, polysaccharides, a matrix protein,and any functional fragment or (chimeric) derivative of the above, aswell as any combinations thereof. For “functional fragment” is hereinmeant any portion of an active agent able to exert itsphysiological/pharmacological activity. For example, a functionalfragment of an antibody could be an Fc region, an Fv region, aFab/F(ab′)/F(ab′)₂ region and so forth.

The bioactive compounds can be added to the scaffold material or to thepolymeric gels by using any suitable method known in the art, such assurface absorption, physical immobilization, e.g., using a phase changeto entrap the substance in the scaffold material, and the like. Forexample, a growth factor can be mixed with a polymeric (e.g. collagen)composition while it is in an aqueous or liquid phase, and after achange in environmental conditions (e.g., pH, temperature, ionconcentration), the liquid gels or solidifies, thereby entrapping thebioactive substance. Alternatively, covalent coupling, e.g. usingalkylating or acylating agents, is used to provide a stable, long-termpresentation of a bioactive substance on the scaffold in a definedconformation. Alternatively, non-covalent adsorption can be used, forexample electrostatic, hydrophobic, dipole-dipole, hydrogen bonding,physisorption and the like. In an additional or alternative embodiment,bioactive molecules can be encapsulated within nano/micro spheres orbeads included within the scaffold and/or a polymeric gel during orafter a manufacturing step. Particularly suitable nano//micro beads arethose described in European Patent Application EP15193284.5,incorporated herein in its entirety by reference.

The scaffold material described herein is useful in the treatment of avariety of diseases, disorders, and defects where surgery or a tissueengineering approach can be a suitable therapeutic solution. This isparticularly true for soft tissues, and the scaffold material maytherefore be utilized in a variety of surgical procedures as well as forcosmetic purposes, and for the treatment or prevention of a plethora ofpathological conditions. The scaffold material of the invention resultsparticularly convenient for treating tissues or organs like agenitourinary tract components, including kidney, bladder, urethra,ureter, vaginal and pelvic floor tissue, a blood vessel (including bigveins and arteries), a muscle, a cartilage, skin, liver, an eye portionsuch as cornea, conjunctiva or sclera, a Central Nervous System (CNS) orPeripheral Nervous System (PNS) component including cerebral ormeningeal tissues, trachea, oesophagus, lungs, stomach, heart,sphincters, pancreas, gut, pharynx or inner ear tissue.

The scaffold is usually transplanted on or close to a target tissue,introduced into or onto a bodily tissue/organ of a subject or a hostusing a variety of methods and tools known in the art, preferably viaminimally invasive surgical devices and procedures.

The term “subject” or “host” as used herein refers to animals,particularly to mammals. For example, mammals contemplated by thepresent invention include humans, primates, pets, domesticated animalssuch as cattle, sheep, pigs, horses, rabbits, rodents and the like.

As used herein, “treatment”, “treating” and the like generally meanobtaining a desired pharmacological and/or physiological effect. Theeffect may be prophylactic in terms of preventing or partiallypreventing a disease, symptom or condition thereof and/or may betherapeutic in terms of a partial or complete cure of a disease,condition, symptom or adverse effect attributed to the disease. The term“treatment” or “treating” as used herein covers any treatment of adisease in an animal, particularly in a mammal, more particularly ahuman, and includes: (a) inhibiting the disease, i.e., arresting itsdevelopment; or (b) relieving the disease, i.e., causing regression ofthe disease and/or its symptoms or conditions such as improvement orremediation of damage. The term “prevention” or “preventing” relates tohampering, blocking or avoid a disease from occurring in a subject whichmay be, for any reason, predisposed to the disease but has not yet beendiagnosed as having it for example based on familial history, healthstatus or age.

Initially, the method for producing the scaffold of the inventionforesees a first fast drying step via a mechanical compression of apolymeric gel, preferably hydrogel, layer (FIG. 1A). It has beenassessed that this first step of mechanical compression allows toachieve a non-uniformly stiff micro/nano structure of the gel and,consequently, of the final scaffold. When working with polymeric gellayers in form of sheets, the mechanical compression is generally doneby applying a rolling or linear compressive force, which must beorthogonal to the planar surface of the sheet. This latter is e.g.placed on a nylon mesh positioned on tissue paper to augment the waterabsorption upon leakage following the compression. When working withpolymeric gel layers already in form of tubes (FIG. 1B), these lattercan be rolled against an adsorbent tissue, with or without e.g. a nylonmesh positioned in between, via e.g. a glass rod or otherwise acylindrical mandrel placed within the lumen of the tube, until no morewater is seen on the tissue paper, for instance by performing 10revolutions of about 30 seconds each on the tissue paper. In someembodiments, an additional filtration step to mechanically suck waterfrom a filter in contact with the lumen can be performed. Tubularpolymeric single layers can be obtained by pouring a polymeric gelprecursor solution into a substantially cylindrical mould comprisingtherein a removable, coaxial elongated support (e.g. a glass rod) of anysuitable diameter so to create a void space, the tube's lumen, uponformation of the polymeric gel (through e.g. gelation of the precursorsolution; FIG. 2).

A key feature of the method according to the invention relies in asecond, slow drying step of the polymeric gel in order to obtain thefinal single layer polymeric scaffold. This step envisages the use oftechniques such as water filtration or preferably air drying, as well ascombinations thereof, for gradually and gently reducing the watercontent of a polymeric gel up to a polymer mass fraction in the finalscaffold of at least 60% w/w. Preferred water filtration means can befor instance dialysis means. In the most preferred embodiment, themethod is characterized in that the second slow drying step is performedby air-drying under gravity for at least 1 hour, preferably for at least4 hours. The air drying process is preferably carried out at roomtemperature (i.e., between about 20° and about 30° C.), and in someembodiments in a controlled atmosphere to keep sterility of the finalscaffold. Once obtained a single layer polymeric scaffold, this can bepossibly circumferentially wrapped around a suitable support of anyshape, such as for instance a removable glass rod, in order to realize atubular Collagen scaffold comprising a lumen therein.

The so-obtained slowly-dried scaffolds feature peculiar characteristicsboth in terms of mechanical values when they are sutured in a surgicalapproach, as well as for what concern tissue regeneration and healing ina host, compared to solutions known in the art. In fact, the scaffoldsof the invention offer an excellent compromise between these two issues,that must contemporary be considered when it comes to the clinics.Current clinically used natural materials sold as off-the shelfsolutions include for instance freeze-dried collagen scaffolds,cross-linked collagen scaffolds and decellularized human and animaltissue-derived scaffolds. These products are designed with a surgicalmind set of being easy to surgically work with. The scaffolds of theinvention, on the other hand, are strong enough to be sutured whileproviding in parallel a more “cellular friendly” matrix, compared to thecommercially available solutions, i.e. surgeon-friendly and stiffscaffolds with poor cellular responses. As a way of example,commercially available tubular acellular materials do not work inurethral replacement of length defects longer than 1 cm; if cells areplaced within these commercial available materials they can achievesuccessful grafts for patients. On the other hand, the presentcollagen-made scaffolds produced according to the methods of theinvention show successful results with 1 cm long or even bigger (e.g., 2cm long) tubular grafts without the addition of any cells. In the sameway, planar scaffolds such as sheets or patches sized at least 1 cm² canbe successfully used in vivo.

In this context, without being bound to this theory, the presentinventors observed that very mechanically stable donor tissues, whichhas been decellularized or devitalized to not elicit immune response inthe graft host patients, are too stiff, or otherwise not enough soft, togive a functional repopulation of the grafted scaffold by host cells.The extreme internal complexity and/or density creates a sort of“extracellular matrix (ECM)-mismatch” between the grafted matrix and theinjured/tissue regeneration site. This mismatch is plausibly the causefor an ill-favoured cell ingrowth, leading to failure of the therapy forthe patients. The scaffold of the invention targets instead a novel“Mechanical and ECM-complexity niche”: it is structurally less complex,less stiff, less dense and therefore easier to be invaded by endogenoushost cells without giving strong inflammatory response, thus favouring abetter and accelerated regeneration profile than other materials, butstill easily suturable. The method of the invention allows to producescaffolds targeting this specific “Mechanical and ECM-complexity niche”.

For instance, in the case of the urinary tract, the scaffold's surfacestiffness favours a better and faster urothelial coverage, which isprobably linked to a better protective environment for sub-urothelialtissue regeneration (like muscle and vascular regeneration). Withoutbeing necessarily bound to this theory, at least one layer in thescaffold shall lay close to native adult tissue in terms of mechanicalproperties for having a better surgical experience, but a majorcomponent should be protected to be soft. This could explain whydecellularized native tissues, having mechanic properties very close tothose of native adult tissues, fail or do not have perfect results;mimicking the mechanics of early development tissues, at least onportions of a scaffold material, could be physiologically more fittingand “better sensed” by a receiving host. In this context, in preferredembodiments of the invention, the stiffer portion of the polymericscaffold should have a retention strength of at least 0.2 N upon suture,preferably of at least 0.4 N, even more preferably with values comprisedbetween about 1.5 and 2 N.

In terms of swelling behaviour, which then also reflects the mechanicalproperties of the scaffold of the invention (water content in polymericscaffolds are known to be related to the mechanical properties, i.e. themore water content the lower mechanical strength), it has beensurprisingly assessed by the inventors that the scaffolds produced withthe present methods, when immersed in an aqueous solution such asPhosphate buffered saline (PBS), re-swell and increase their watercontent only to up of a maximum of 12% w/w at steady state. Compared toother manufacturing techniques, this has important advantages in termsof functional and mechanical properties of the final scaffold: byapplying a fast mechanical compression to a first polymeric gel followedby a slow drying step, a scaffold ready to be sutured can be easilyachieved. In comparison, a freeze-dried collagen gel swells and increaseits water content to more than 20% w/w (FIG. 5B). Mechanicallycompressed collagen gel swells and increases its water content only by0.7%. Substantially, the developed method permits to better control there-swelling behaviour of the obtained scaffold, and then ultimately themechanical and functional properties ideal for surgery.

This property represents moreover a huge advantage for what concerns thestocking and the packaging of the scaffold of the invention, since itcould be provided to a surgeon as an off-the-shelf, ready-to-usesolution. In fact, it can be imagined a polymeric scaffold of theinvention prepared and packaged in a liquid/aqueous solution alreadyadapted for use in an in vivo approach, without any change in itsproperties at steady state even when hydrated. The “re-swelling steadystate” is defined as the state in terms of water content of thepolymeric scaffold produced with the methods of the invention in whichthe mechanical properties and the water content of the said scaffold donot change anymore.

With the aim of accentuating the mechanical differences between twodifferent areas of a polymeric scaffold of the invention, the inventorsdeveloped a very ingenious and elegant method for producingnon-uniformly stiff scaffolds, said method being characterized in thatit comprises the following steps:

a) providing a single layer of a mechanically compressed polymericmaterial gel wrapped around a support;

b) dipping a portion of the support-wrapped polymeric gel layer into anaqueous solution while leaving another portion air-drying under gravityfor a certain amount of time;

c) flipping the polymeric gel layer portion previously dipped into theaqueous solution onto the air-dried portion; and

d) optionally air-drying under gravity both layers for a certain amountof time

wherein at least one layer of the so obtained polymeric scaffold has apolymer mass fraction of at least 60% w/w. This method, renamed by theinventors “the sock-technique” and shown in FIGS. 4 and 5, can beapplied, mutatis mutandis, to polymeric gels formed by e.g. pouring apolymeric gel precursor solution into a substantially cylindrical mouldcomprising therein a coaxial elongated support for creating a tubularsingle layer of polymeric gel, gelling it and mechanically compress it.

An additional object of the invention relates to a kit comprising acontainer, an aqueous solution and the polymeric scaffold of theinvention. The three components can be provided separated within thekit, or the container can already include the liquid solution and alsopossibly the polymeric scaffold in a re-swelled, steady-state form.Preferably, the entire kit is sterilized or sterilisable by electrodebeam, gamma rays or X-rays radiations.

EXAMPLES

The present, non-limiting examples present one embodiment of thefabrication method of a non-uniformly stiff tubular collagen scaffoldaccording to the invention. The method can be easily adapted to makesheet and dome-shaped structures, and dimensions and shape of thecollagen scaffold depend only of the used mold. For the fabrication of atubular non-uniformly stiff collagen scaffold with an inner lumen of 3mm, 8.5 mL of liquid acido-soluble bovine collagen solution (5 mg/mL,Symatese, France) are mixed with 0.8 mL 10× MEM (Invitrogen) andneutralized by the addition of approximately 1.85 mL of 0.1M NaOH. Theexact volume of NaOH to be added is determined by observing a colourchange from yellow to pink. The neutralized collagen solution is pouredinto a tubular shaped, stainless steel mould (length 70 mm, innerdiameter 12 mm). The lumen of the tubular collagen scaffold is formed bya glass stick (diameter 3 mm) placed inside the tubular mould. Collagengelation takes place at room temperature for 15 min. After gelation, thecollagen tube is removed from the mould but kept on the glass stick.This gelated collagen tube is the starting matrix to produce non-uniform(denser external walls and a less dense internal side) tubular collagenscaffolds.

The tubular collagen gel is mechanically compressed by rolling on aroundits own axis, to ensure unequal compression and drainage of water withinthe final scaffold. The collagen gel is rolled over a nylon mesh placedon five layers of double-layer tissue paper (Weita, Arlesheim,Switzerland) until no more water is coming out. Subsequently, a furtherdrying step is performed by exposing the collagen tube to air at roomtemperature for 1 to 24 hours. The desired dryness corresponds to0.5-40% of the initial water content of the gel, preferably 18-22% ofwater. The weight of the collagen gel after gelation represents 100% ofwater. The percentage of water in each dried collagen scaffold iscalculated by taking the difference of the collagen gel weight beforeand after drying:

% of water=1−[(weight before drying−weight after drying)/weight beforedrying]

Once the collagen tube has reached the desired dryness, the tube isplaced in PBS (pH 7.4) for 3 days in order to re-swell it to asteady-state water content. The water content of the final collagenscaffold is in the range of 5 to 48%, preferably 24-30% of watercontent.

For the fabrication of a multi-layered collagen tube through the “socktechnique”, half of a tube dried with a first rolling step as previouslydescribed is placed into a rehydration bath (PBS, pH 7.4) for 1 to 60minutes, preferably for 30 minutes. The rehydrated half tube is thenpulled over the dried tube portion or the dry portion is pushed underthe rehydrated part for creating a multi-layered tube. Multi-layeredtubes (with more than two layers) can be obtained by iterating the abovedescribed steps.

Ultimate tensile strength (UTS) and Young's modulus were determinedunder stress strain conditions using an Instron tensile machine(Norwood, Mass., USA), applying a 250 N load with a strain rate of 1mm/min until the break point. Tubular samples of 5 mm length was pulledfrom the lumen side in one direction using L-shaped hookes placed in thelumen. Young's modulus values were obtained from the slope of theinitial linear region of the stress-strain curves, while the UTS wereobtained from the maximum tensile stress recorded.

To determine the collagen density and porosity in the scaffolds,scanning electron microscopy (SEM) was performed. The samples were fixedwith 1% tannic acid and 1.25% glutaraldehyde, then washed with 0.1 Mcacodylate, and dehydrated in increasing ethanol concentrations prior tocritical point drying (CPD). They were coated with gold/palladium andimaged at a voltage of 10 kV using a scanning electron microscope (SEM,XLF30, Philips). To further visualize the collagen density and porositywithin the scaffolds, formalin fixed (4% PFA) and paraffin embedded cutsections (8 μm thick) were stained with anti-collagen typel antibodies(1:250 dilution, rabbit anti-collagen type 1, Abcam, Switzerland).Corresponding secondary AlexaFluor antibodies (Abcam, Switzerland) wereused to visualize the type 1 collagen distribution under a LSM700microscope (Zeiss, Germany).

To determine the in vitro suture capacity, anastomosis of two producedcollagen tubes was performed using Vicryl 6.0® sutures. Furthermore, todetermine the in vivo suture capacity of the produced tubes, at least 9rabbits had a urethral replacement surgery, where a double end-to-endanastomoses to the native urethra were performed with interruptedsutures (Vicryl 6.0®). Rabbits were sacrificed after 1 and 3 months.Thereafter, rabbit penises were harvested and fixed in 4% formalin,embedded in paraffin, and sectioned in 8 μm thick for Hematoxylin-eosinstaining. Images were taken with a Leica DM5500 microscope (Germany).

To understand different cell type behaviour (e.g. urothelial and smoothmuscle cells) on various collagen-based scaffold with differentstiffness and collagen fiber densities, human urothelial cells that werelentiviral-transduced with GFP were seeded onto these collagen scaffoldsand cell morphology was observed under a fluorescence stereomicroscope(Leica). The human cells were placed in a fibrin drop to control thespatial distribution of cells on the collagen, and the cells at varioustime points were then imaged to verify the cells fate choices.

Results

First of all, mechanical compression with air-drying and only air-dryingyields different materials, as seen in FIG. 5A it also takes longer timeto reach different dry states for only air-dried material compared tomechanical compression with air-drying. A fast compression yields lessre-swelling compared to the other method of slow drying. As shown inFIG. 5B, the collagen scaffolds dried with either freeze-drying (FD)air-drying (AD) or mechanical compression with air drying (M-AD)re-swell when they are placed in hydrating conditions (FD dry: 1.3% toFD re-swell: 20%, AD dry: 1.5 to AD re-swell: 9.5%, M-AD dry: 0.8% toM-AD re-swell 5.2% water content). However, mechanically compressed (M)have a very low swelling capacity when placed in hydrating conditions (Mdry: 47.6% to M re-swell: 48.3% water content). Furthermore, themechanical compressed plus air-dried collagen scaffold's re-swelling canbe controlled by its dry state, as shown in FIG. 5C. There is a linearrelationship between the dry state and the re-swelling state of theproduced collagen scaffolds. The re-swelling in hydration (in PBS-bath)is a rapid process where most of the swelling has occurred within thefirst 30 minutes (FIG. 6). A more stable re-swelling steady-state isseen after 72 h in a hydrating environment. The mechanical properties,Young's modulus and UTS of the collagen scaffolds with different watercontents follow an exponential curve (FIGS. 7A and 7B). As shown inFIGS. 7C and 7D, the collagen tube dried and re-swelled to a 9.5% watercontent is sutured with ease into rabbits and goat urethra's (goaturethra is shown in FIG. 7D). Also the re-swelled collagen scaffold of45% water content is sutured to rabbit and dog urethra (dog urethrasurgery is seen in FIG. 7E). Classifying what is suture compatible isdifficult since it is highly dependent on the surgeon, and can only bedone on an empirical basis. However, starting from the mechanicalproperties recorded and surgeons opinion when handling the collagengrafts of different water content, a material that has a dry state of atleast one collagen layer of no more than 40% w/w water content (i.e., atleast 60% w/w collagen mass fraction) was classified as suturecompatible directly after it has been re-swelled to steady-state inhydrating condition, when produced from a mammalian collagen precursorliquid solution with range of collagen concentration comprised between2.1 and 40 mg/mL.

The high functional regeneration potential of the scaffold of theinvention is highlighted in a surgical setting where collagen tubesproduced according to the methods of the invention were implanted andsutured in a rabbit urethra. The collagen scaffolds stay in the suturedarea as shown in a histology section of 1 month after surgery (FIG. 7E),black arrows show non-degraded collagen from remaining parts of thecollagen scaffolds). After 3 months, no more of the implanted collagencan be visualized by histology, since the rabbits' own cells haveremodelled the collagen and populated the area (FIG. 7F).

A crucial observation done by the inventors is that a slow-dryingprocedure creates an uniform and homogenous collagen architecture atmicro- and nanoscale (FIGS. 8B and 8D). While a freeze-dried collagenstructure yields an unorganized non-uniformed architecture with randomdistribution of holes (diameter of 50-150 μm) (FIGS. 8C and 8F) at microand nanoscale. In comparison, if one combine a first mechanicalcompression followed by a slow-drying step, this yields to a non-uniformbut organized heterogeneous collagen architecture at micro- andnanoscale, where the surface that is in close contact with themechanical force gets a more dense structure as compared to the internalside (FIGS. 8A and 8C). In FIG. 9, the same feature is shown byutilizing immunohistochemistry to visualize collagen type 1.

Many studies report that material stiffness, fiber density and porosityinfluence cell fate choices, and that different cell types behavedifferently by changing the material properties. FIG. 10 depicts thebehaviour of urothelial cells in response to collagen matrices thatdiffer in stiffness. In a less dense collagen matrix, the urothelialcells are less in contact with neighbours and migrate away from theirstarting position. On the other hand, urothelial cells seeded on adenser scaffold stay more in contact with neighbour cells and migrateless.

To further increase mechanical stability and the potential for creatingnon-uniformly stiff materials, with an option of spatially distributegrowth factors and reagents to modify the collagen at various positionin a scaffold, a further method, shown in FIG. 3, was developed. Itdescribes a “sock-technique”, taking advantage of the drying andswelling behaviour of the collagen described before. Variousmanipulations can be performed in order to get a structure with singlelayers or multiple layers, by pulling or pushing over a wetted structureon top of the other. In FIG. 4 it is shown the technique and the finalstructure of a double-layered and single-layered tube, wherein theformer is clearly more stable since it can hold its own weight withouthaving the lumen collapsing. Furthermore, tubes produced in this way canalso be opened up by e.g. blade cut, resulting in single- ormulti-layered sheets, which can be used as patches for surgical use. Adouble-layered sheet, originating from an opened double-layered tubeusing a scissor, is shown in FIG. 4H.

Additionally, the bioactivity of the non-uniform collagen scaffoldscould be enhanced by the addition of bioactive molecules to therehydration bath using one of the known platforms to efficiently bindbioactive molecules to collagen resulting in a controlled release fromthe collagen scaffold. Collagen scaffolds prepared according to thedescribed method are highly useful for soft tissue engineering orsurgical applications of organs/tissues such as genitourinary system(i.e. for bladder augmentation, vaginal and pelvic floor reconstruction,ureter and/or urethra reconstruction), for vascular tissue engineering,for the reparation of the oesophagus or as patches for hernia repair.

1-25. (canceled)
 26. A polymeric scaffold having a non-uniform densityfor use in tissue engineering, diagnostics, or surgical procedures,wherein the polymeric scaffold has a polymer mass fraction of at least60% w/w of a total scaffold weight, and wherein the polymeric scaffoldhas a density at a surface that is denser as compared to a density at acore.
 27. The polymeric scaffold of claim 26, wherein a pore diameter atthe core is between 0.2 μm and 200 μm.
 28. The polymeric scaffold ofclaim 26, wherein an average pore diameter at the surface is less than0.2 μm.
 29. The polymeric scaffold of claim 26, wherein a Young'smodulus at the core is between 0.1 kPa and about 20 kPa.
 30. Thepolymeric scaffold of claim 26, wherein a Young's modulus at the surfaceis between 5 kPa and 1500 kPa.
 31. The polymeric scaffold of claim 26,wherein the polymeric scaffold is suturable.
 32. The polymeric scaffoldof claim 26, wherein the polymeric scaffold is at least 1 cm long. 33.The polymeric scaffold of claim 26, wherein the polymeric scaffoldincludes a bioactive agent.
 34. The polymeric scaffold of claim 26,wherein the polymeric scaffold is cell-free.
 35. The polymeric scaffoldof claim 26, wherein the polymeric scaffold includes cells.
 36. Thepolymeric scaffold of claim 26, wherein the polymeric scaffold isimmersed in an aqueous solution.
 37. The polymeric scaffold of claim 26,wherein the polymeric scaffold does not include a crosslinking agent.38. The polymeric scaffold of claim 26, wherein the polymeric scaffoldis configured to reswell in an aqueous solution by more than 4% w/w toreach a steady-state water content.
 39. The polymeric scaffold of claim38, wherein the polymeric scaffold is configured to reswell in a rangeof 5% to 48% w/w to reach a steady-state water content.
 40. Thepolymeric scaffold of claim 26, comprising one or more polymericmaterial layers.
 41. The polymeric scaffold of claim 26, comprising oneor more polymeric tubes.
 42. The polymeric scaffold of claim 26, whereina polymeric material of the polymeric scaffold includes gelatin,elastin, collagen, agar/agarose, chitosan, fibrin, methyl cellulose,hyaluronic acid, proteoglycans, polyamino-acids or derivative thereof,polysaccharides or derivatives thereof, glycosaminoglycans orderivatives thereof, or any combination of the foregoing.
 43. An implantkit comprising: a container; an aqueous solution located inside thecontainer; a polymeric scaffold immersed by the aqueous solution insidethe container, wherein the polymeric scaffold has a polymer massfraction of at least 60% w/w of a total scaffold weight, and thepolymeric scaffold has a density at a surface that is denser as comparedto a density at a core.
 44. A graft made of a polymeric scaffold havinga non-uniform density for surgical applications, wherein the polymericscaffold has a polymer mass fraction of at least 60% w/w of a totalscaffold weight, and wherein the polymeric scaffold has a density at asurface that is denser as compared to a density at a core.
 45. Thepolymeric scaffold of claim 44, wherein the polymeric scaffold isconfigured to reswell in an aqueous solution by more than 4% w/w toreach a steady-state water content.